In magnetic resonance imaging, the spins of specific nuclei (usually hydrogen nuclei) in a tissue are excited by radiofrequency (RF) pulses in the presence of an applied static magnetic field in a selected direction, the magnitude of which is made to spatially vary in a defined time sequence. The precessional frequencies of the excited spins vary in relation to the magnitude of the applied magnetic field and thereby produce a precessional signal from which the spatial locations of the spins can be derived. By applying one or more excitation RF pulses and a specific sequence of linear spatial variations in the applied magnetic field, referred to as gradient pulses, the resulting precessional signal can be interpreted as a carrier waveform amplitude modulated by the Fourier transform of the spatial distribution of spin density in a selected portion of the tissue. The carrier waveform in this case is a complex sinusoid at the spin resonance frequency with no gradient applied (i.e., the Larmor frequency of the spin species). Transformation from the spatial frequency domain, referred to as k-space, to the image position domain can be accomplished by inverse Fourier transforming the k-space signal which is generated after demodulation of the precessional signal. The k-space signal is thereby transformed to a spin density function in position space which can be used to generate an image where the intensity of an image pixel varies in accordance with the magnitude of the spin density function at the pixel location. In order to image a selected volume of interest (VOI) in the body, an MRI data set is acquired which is made up of a plurality of slices derived from a two-dimensional (2D) spin density function or a plurality of slabs derived from a three-dimensional (3D) spin density function. As the term is used herein, “image” should be taken to mean either an actual visual representation or the data from which such a representation could be rendered. Similarly, a “pixel” or “voxel” should be taken to mean either a discrete element of an actual 2D or 3D visual representation, respectively, or the corresponding element of a 2D or 3D object from which such a representation could be rendered.
The time sequence of RF excitation and gradient pulses may be manipulated so that the spin density function derived from the k-space signal is dependent upon other parameters in addition to spin density, such as the spin-lattice relaxation time constant T1 or the spin-spin relaxation time constant T2. The time constant T1 relates to the time required for spins to recover longitudinal magnetization after an excitation pulse, the longitudinal magnetization being necessary for the generation of an FID signal following an excitation pulse. A pulse sequence may be designed so that spins with a shorter T1 are weighted more heavily in the spin density function, and a so-called T1 weighted image may be derived from such a spin density function. The time-of-flight (TOF) method of imaging blood flow in tissue involves the use of repeated excitation pulses timed so that blood flowing from an unexcited region into the region excited by the pulses has a greater longitudinal magnetization than the stationary tissue in the excited region. The moving blood thus mimics a tissue with a short T1 and produces an enhanced spin signal. TOF imaging may be used to selectively image blood vessels owing to the moving blood contained within the vessels.
Blood flow may be imaged and quantified by another technique, phase contrast magnetic resonance (PCMR). The k-space signal from the excited spins is a complex signal in which the real and imaginary components modulate the carrier waveform in phase quadrature. Ideally, inverse Fourier transformation of the k-space signal results in a purely real spin density function. Certain artifacts may cause the spin density function to have both real and imaginary parts, but this problem can be circumvented in normal imaging by varying the image pixel or voxel intensity in accordance with the magnitude of the spin density function to create a so-called magnitude image. In PCMR, on the other hand, a bipolar gradient pulse is used to cause flowing spins to acquire a phase which is proportional to the velocity of the spins in the direction of the gradient. After such phase-encoding of velocity, the phase can be extracted from the spin density function to measure the magnitude of blood flow. The extracted phase can also be used to construct an image where the pixel or voxel intensity varies with the phase of the spin density function at the location of the pixel or voxel, called a phase image. A phase image derived from a k-space signal derived after application of an appropriate through-plane bipolar gradient pulse can thus provide a visual representation of the magnitude of blood flow through the plane of the image.